DEVELOPMENT OF A SCAFFOLD INCORPORATING

ZINC-OXIDE NANO-PARTICLES FOR CARTILAGE TISSUE

ENGINEERING UNDER PHYSIOLOGICAL CONDITIONS

ERAJ HUMAYUN MIRZA

FACULTY OF ENGINEERING

UNIVERSITY OF MALAYA

KUALA LUMPUR

2016

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DEVELOPMENT OF A SCAFFOLD INCORPORATING

ZINC-OXIDE NANO-PARTICLES FOR CARTILAGE

TISSUE ENGINEERING UNDER PHYSIOLOGICAL

CONDITIONS

ERAJ HUMAYUN MIRZA

THESIS SUBMITTED IN FULFILMENT OF THE

REQUIREMENTS FOR THE DEGREE OF DOCTOR OF

PHILOSOPHY

FACULTY OF ENGINEERING

UNIVERSITY OF MALAYA

KUALA LUMPUR

2016

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UNIVERSITY OF MALAYA

ORIGINAL LITERARY WORK DECLARATION

Name of Candidate: Eraj Humayun Mirza

Matric No: KHA110058

Name of Degree: Doctor of Philosophy

Title of Thesis: The Development of a Scaffold incorporating Zinc-Oxide Nano-particles

for Cartilage Tissue Engineering under Physiological Conditions

Field of Study: Biomaterials and Tissue Engineering

I do solemnly and sincerely declare that:

(1) I am the sole author/writer of this Work;

(2) This Work is original;

(3) Any use of any work in which copyright exists was done by way of fair dealing and

for permitted purposes and any excerpt or extract from, or reference to or reproduction of

any copyright work has been disclosed expressly and sufficiently and the title of the Work

and its authorship have been acknowledged in this Work;

(4) I do not have any actual knowledge nor do I ought reasonably to know that the making

of this work constitutes an infringement of any copyright work;

(5) I hereby assign all and every rights in the copyright to this Work to the University of

Malaya (“UM”), who henceforth shall be owner of the copyright in this Work and that

any reproduction or use in any form or by any means whatsoever is prohibited without

the written consent of UM having been first had and obtained;

(6) I am fully aware that if in the course of making this Work I have infringed any

copyright whether intentionally or otherwise, I may be subject to legal action or any

other action as may be determined by UM.

Candidate’s Signature Date: 22nd August 2016

Subscribed and solemnly declared before,

Witness’s Signature Date: 22nd August 2016

Name:

Designation:

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Abstract

Cartilage tissue is among the most complex of all in the human body. Degeneration of

cartilage, lack of repair and various traumatic and pathological conditions among

individuals increase joint pain and disability. Researchers have been trying to repair

cartilage damage for decades, but have failed to achieve an optimal repair strategy.

Cartilage tissue is adapted to low oxygen environments and this condition appears to be

a key factor in the growth, regulation, and survival of chondrocytes; which are the only

cells present in cartilage tissue. The present thesis defines the fabrication and

characterisation of an antibacterial biomaterial for cartilage tissue repair. Furthermore,

this thesis promotes the use of Zinc Oxide (ZnO) nanoparticles (NP) for better growth of

cartilage cells and survival of cartilage tissue as a whole.

Composite scaffolds and thin films of polyoctanediol citrate (POC) polyester elastomer,

with varying concentrations (1wt%, 3wt% and 5wt%) of ZnO were fabricated by a

solvent-casting/particulate-leaching and mould casting technique respectively. It was

observed that material properties can be successfully controlled by simple variation of

NP concentration within the composite. The ion release kinetics from ZnO-POC scaffolds

are strongly dependent on NP concentration and degradation of pure POC matrix. All the

composite scaffolds showed strong antibacterial characteristics. However, cell culture

studies demonstrated that 1% ZnO incorporation in POC polymer is the optimal

concentration for chondrocyte cells.

Moreover, the effect of 1% ZnONP on chondrocyte proliferation and matrix synthesis

cultured under normoxia (21% O2) and hypoxia (5% O2) demonstrated upregulation of

chondrocyte proliferation and sulphated glycosaminoglycan (S-GAG) in hypoxic culture.

A synergistic effect of oxygen concentration and 1% ZnONP in up-regulation of anabolic

gene expression (Type II collagen (COL2A1) and aggrecan (ACAN)), and a down

regulation of catabolic (MMP-13) gene expression was observed. Furthermore,

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production of transcription factor hypoxia-inducible factor 1A (HIF-1A) in response to

hypoxic condition to regulate chondrocyte survival under hypoxia was not affected by the

presence of 1% ZnONP. Lastly this thesis discusses the physiological adaptations of

cartilage tissue in a dynamic mechanical loading and hypoxic condition to yield benefits

of combined bio-factors for cartilage tissue engineering applications. The results indicate

that the combination of dynamic loading is compatible with the nanoparticle addition.

Furthermore, dynamic loading suppresses MMP-13, and increase the expression of

COL2A1 and ACAN with an increase in cell viability, and promotion of rounded cell

morphology (a phenotypic marker).

It was concluded that POC-ZnONP scaffolds are of major importance in the development

of multifunctional scaffolds based on biodegradable polyesters for cartilage tissue

engineering and presence of 1% ZnONP appears to preserve homeostasis of cartilage in

its hypoxic environment. While 1% ZnONP should be considered for beneficial

incorporation into 3D hypoxic culture systems in the presence of mechanical stimulation.

Further studies must focus on determining the use of 1% ZnONP in different polymers

for use in various tissue engineering applications.

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Abstrak

Tisu rawan adalah antara tisu yang paling kompleks di dalam badan manusia.

Kemerosotan rawan, kurang pembaikan dan pelbagai trauma dan keadaan patologi di

kalangan individu meningkatkan kesakitan pada sendi dan kecacatan. Penyelidik telah

berusaha untuk memperbaiki kerosakan rawan selama bertahun-tahun, tetapi telah gagal

mencapai strategi optimum pembaikan. Tisu rawan disuaidirikan kepada persekitaran

rendah oksigen dan keadaan ini menjadi satu faktor kepada pertumbuhan, peraturan dan

kebertahanan kondrosit; iaitu satu-satunya jenis sel yang ada di dalam tisu rawan. Tesis

ini menentukan fabrikasi dan pencirian biobahan antibakteria untuk pembaikan tisu

rawan. Tambahan pula, tesis ini menggalakkan penggunaan nano-partikel (NP) Zink

Oksida (ZnO) untuk pertumbuhan yang lebih baik bagi sel-sel rawan dan kebertahanan

tisu rawan secara keseluruhannya.

Perancah komposit dan filem nipis daripada elastomer polyester polyoctanediol sitrat

(POC), dengan kepekatan (1wt%, 3wt% and 5wt%) ZnO yang berbeza telah

difabrikasikan masing –masing dengan menggunakan teknik pelarut-pemutus/larutresap-

zarahan dan pemutus acuan. Daripada pemerhatian, didapati sifat-sifat material telah

berjaya terkawal oleh perubahan mudah kepekatan NP di dalam komposit itu. Kinetik

pembebasan ion daripada perancah ZnO-POC bergantung secara kuat kepada kepekatan

NP dan degradasi matriks tulen POC. Semua perancah komposit telah menunjukkan sifat-

sifat antibakteria yang kuat. Walau bagaimanapun, kajian sel kultur mendapati gabungan

1% ZnO di dalam polimer POC adalah kepekatan yang optimum untuk sel-sel kondrosit.

Tambahan lagi, kesan 1% ZnONP terhadap pertumbuhan kondrosit dan sintesis matriks

yang dikultur di bawah keadaan normoksia (21% O2) dan hipoksia (5% O2) telah

menunjukkan pengawalan-atas pertumbuhan kondrosit dan sulfat-glycosaminoglycan (S-

GAG) di dalam kultur hipoksia. Satu kesan sinergi kepekatan oksigen dan 1% ZnONP di

dalam penyataan gen pengawalan-atas anabolic (Kolagen Type II (COL2A1) dan

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aggrecan (ACAN)), dan penyataan gen pengawalan-bawah katabolic (MMP-13) telah

diperhatikan. Tambahan pula, pengeluaran hipoksia transkripsi faktor-dorongan factor

1A (HIF-1A) di dalam tindakbalasnya terhadap keadaan hipoksia untuk mengawal

kewujudan kondrosit tidak dipengaruhi dengan kehadiran 1% ZnONP. Akhir sekali, tesis

ini membincangkan tentang penyesuaian fisiologi tisu rawan di dalam pemuatan

mekanikal dinamik dan keadaan hipoksia untuk menghasilkan manfaat biofaktor yang

digabungkan untuk penggunaan kejuruteraan tisu rawan. Keputusan menunjukkan

gabungan pemuatan dinamik adalah serasi dengan penambahan nano-partikel. Tambahan

pula, pemuatan dinamik menyekat MMP-13 dan meningkatkan penyataan COL2A1 dan

ACAN dengan peningkatan pertumbuhan sel dan morfologi sel bulat (satu penanda

fenotip).

Ini dapat disimpulkan bahawa perancah POC-ZnONP sangat penting di dalam

penambahbaikan pelbagai fungsi perancah berasaskan polyester biodegradasi untuk

kejuruteraan tisu rawan dan kehadiran 1% ZnONP telah memelihara homestasis rawan di

dalam persekitaran hipoksia. Manakala 1% ZnONP perlu dipertimbangkan untuk

gabungan bermanfaat kepada system 3D kultur hipoksia dengan kehadiran rangsangan

mekanikal. Kajian lanjut fokus kepada penentuan penggunaan 1% ZnONP di dalam

polimer yang berbeza untuk kegunaan pelbagai aplikasi kejuruteraan tisu.

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I dedicate this thesis to Ammi and Baba

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ACKNOWLEDGEMENT

First of all I would like to praise ALLAH who blessed me with the strength to accomplish

my dream.

I would like to explicitly thank my supervisors Dr. Belinda Pingguan – Murphy and Dr.

Wan Mohd Azhar bin Wan Ibrahim, for their unconditional support, vast knowledge,

precious time and critical feedback. I am especially grateful to Dr. Belinda Pingguan–

Murphy for her patience with me and her critical review which allowed me to enhance

the quality of my articles and my thesis. Moreover, I sincerely acknowledge my ustaad,

Dr. Ali Moradi, from whom I have learnt not only cell culture techniques and scaffold

fabrication but also to help everyone. I would like to extend my appreciation to my

friends, my lab mates and my colleagues; Adel Al-Halawani, Yasser Al-Saffar, Haris

Akram, Salfarina Ezrina, Kar Wey, Poon Chi-Tat, Janu Ru, Dr. Chong Pan-Pan, Dr.

Eleanor Parker, Syed Shahabuddin, Shahid Mehmood, Kashan Pirzada, Forough

Ataollahi, Adel Dalilottojari and all those who supported me knowing or unknowingly.

I like to thank Dr. Wan Safwani and Dr. Farina to provide me with their valuable

suggestions during candidature defence.

I admire the support from Dr. Adib, Dr. Ahmad Khairi, Liyana bint Abu, Kakak Enas

naeem, Mr. Hanafi, Mr Adhli and all the staff at biomedical engineering administration.

I would also like to thank all the laboratory technical staff throughout University of

Malaya.

I would like to acknowledge the scholarship from MOHE without which I would have

never had enough financial support to continue my PhD. Furthermore, I would like to

thank University of Malaya for providing IPPP and HIR grant to fulfil my research needs.

I would like to thank Ammi and Baba for their unconditional love, support and long

conversations to calm down my anxiety. A sincere thanks to my brothers. My gratitude

and love to my wife who loves me and care for me.

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TABLE OF CONTENTS

LIST OF FIGURES ……………………………………………………… xv

LIST OF TABLES ………………………………………………………. xx

LIST OF ABBREVIATIONS …………………………………………… xxi

CHAPTER 1: INTRODUCTION ………………………………………..

1.1 Thesis Overview ………………………………………………………

1.1 Background ……………………………………………………………

1.2 Problem Statement …………………………………………………….

1.3 Objectives ……………………………………………………………...

1.4 Hypotheses …………………………………………………………….

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CHAPTER 2: LITERATURE REVIEW ……………………………….

2.1 Introduction ……………………………………………………………

2.2 Cartilage ……………………………………………………………….

2.2.1 Composition of articular cartilage ………………………….

2.2.1.1 Chondrocytes …………………………………………

2.2.1.2 Extracellular Matrix ………………………………….

2.2.1.3 Water …………………………………………………

2.2.1.4 Collagens ……………………………………………..

2.2.1.5 Proteoglycans ………………………………………...

2.2.1.6 Non-Collagenous proteins and Glycoproteins ……….

2.2.2 Zonal arrangement in articular cartilage ……………………….

2.3 Chondrogenesis and chondrocyte differentiation ……………………...

2.4 Articular Cartilage Disease and Injury ………………………………...

2.5 Treatment Options for Cartilage Injury ………………………………..

2.5.1 Marrow Stimulation Technique ………………………………..

2.5.2 Osteochondral Grafting ………………………………………...

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2.5.3 Autologous Chondrocyte Transplantation ……………………..

2.6 Cartilage Tissue Engineering ………………………………………….

2.6.1 Cell Sources ……………………………………………………

2.6.2 Cell Signalling Strategies ………………………………………

2.6.2.1 Mechanical stimulus ………………………………….

2.6.2.2 Oxygen Tension/Hypoxia ……………………………

2.6.3 Tissue Engineering Scaffolds ………………………………….

2.7 Poly Octanediol Citrate (POC) ………………………………………..

2.7.1 POC, its structure, synthesis, and form ………………………...

2.7.2 POC as a biomaterial …………………………………………..

2.7.3 POC in cartilage tissue engineering ……………………………

2.8 Infection and Anti-bacterial scaffolds …………………………………

2.9 Antibacterial activity of ZnO ………………………………………….

2.10 Motivation for current study …………………………………………..

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CHAPTER 3: MATERIALS AND METHODS ………………………...

3.1 Introduction ……………………………………………………………

3.2 Preparation of POC pre-polymer ……………………………………...

3.3 Fabrication of POC and ZnO-POC composites ……………………….

3.4 Chondrocyte Isolation …………………………………………………

3.4.1 Chondrocyte Culture Medium …………………………………

3.4.2 Cartilage ECM digestion ………………………………………

3.4.2.1 Protease preparation ………………………………….

3.4.2.2 Collagenous preparation ……………………………...

3.4.3 Cell Viability …………………………………………………...

3.5 Cell proliferation assay ………………………………………………..

3.6 DNA Biochemical Assay ……………………………………………...

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3.6.1 Lysis of BACs ………………………………………………….

3.6.2 DNA quantification …………………………………………….

3.6.2.1 DNA Standard ………………………………………..

3.7 Sulphated Glycosaminoglycan (S-GAG) quantification assays ………

3.8 Quantitative Polymerase Chain Reaction (q-PCR) ……………………

3.8.1 RNA Isolation ………………………………………………….

3.8.2 cDNA synthesis ………………………………………………..

3.8.3 Gene Expression ……………………………………………….

3.9 Scanning Electron Microscopy (SEM) for cell adhesion ……………...

3.10 Live/Dead Assay ……………………………………………………..

3.11 Statistical Analysis …………………………………………………

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CHAPTER 4: CHARACTERISATION OF POLYOCTANEDIOL

CITRATE-ZINC OXIDE NANO-COMPOSITE SCAFFOLD ………...

4.1 Introduction …………………………………………………………....

4.2 Materials and Methods ………………………………………………...

4.2.1 Surface and elemental properties ………………………………

4.2.2 Zinc oxide nanoparticles distribution within POC …………….

4.2.3 Surface wettability test …………………………………………

4.2.4 Mechanical property and porosity ……………………………..

4.2.5 Swelling test ……………………………………………………

4.2.6 Fourier transform infrared spectroscopy (FTIR) analysis ……..

4.2.7 Thermal analysis ……………………………………………….

4.2.8 Degradation analysis …………………………………………...

4.2.9 ZnO NPs release kinetics in physiological conditions …………

4.2.10 Anti-bacterial properties ……………………………………….

4.2.11 In vitro tests with chondrocyte cell culture …………………….

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4.3 Results …………………………………………………………………

4.3.1 Surface morphology ……………………………………………

4.3.2 Influence of ZnO-NPs on relative hydrophilicity ……………...

4.3.3 Physical properties of scaffolds ………………………………..

4.3.4 Swelling and cross-linking density …………………………….

4.3.5 Chemical composition …………………………………………

4.3.6 Thermal stability ……………………………………………….

4.3.7 Degradation …………………………………………………….

4.3.8 ZnO release kinetics ……………………………………………

4.3.9 Anti-bacterial properties of ZnO-POC scaffolds ………………

4.3.10 Cytotoxic effects of ZnO NPs on primary chondrocytes ………

4.4 Discussion ……………………………………………………………..

4.4.1 Surface morphology ……………………………………………

4.4.2 Influence of ZnO-NPs on relative hydrophilicity ……………...

4.4.3 Physical properties of ZnO-POC scaffolds: porosity and

elasticity ……………………………………………………......

4.4.4 Swelling and cross-linking density …………………………….

4.4.5 Chemical composition …………………………………………

4.4.6 Thermal stability ……………………………………………….

4.4.7 Degradation …………………………………………………….

4.4.8 ZnO release kinetics ……………………………………………

4.4.9 Anti-bacterial properties of ZnO-POC composite scaffolds …...

4.4.10 Cytotoxic effects of ZnO NPs on primary chondrocytes ………

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CHAPTER 5: EFFECT OF HYPOXIA ON BOVINE ARTICULAR

CHONDROCYTES ……………………………………………………….

5.1 Introduction ……………………………………………………………

5.2 Materials and Methods ………………………………………………...

5.2.1 Experimental Design …………………………………………...

5.2.2 Cell Proliferation Assays ………………………………………

5.2.3 Sulphated glycosaminoglycan (S-GAG) quantification

assays …………………………………………………………..

5.2.4 Quantitative Polymerase Chain Reaction (q-PCR) …………….

5.2.5 Statistical Analysis ……………………………………………..

5.3 Results …………………………………………………………………

5.3.1 Cell proliferation ……………………………………………….

5.3.2 Proteoglycan Synthesis ………………………………………...

5.3.3 Gene expression ………………………………………………..

5.4 Discussion …………………………………………………………......

5.4.1 Cell proliferation ……………………………………………….

5.4.2 Proteoglycan Synthesis ………………………………………...

5.4.3 Gene expression ………………………………………………..

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CHAPTER 6: EFFECT OF DYNAMIC MECHANICAL

COMPRESSION ON BOVINE ARTICULAR CHONDROCYTES ….

6.1 Introduction ……………………………………………………………

6.2 Materials and Methods ………………………………………………...

6.2.1 Experimental Design …………………………………………...

6.2.2 Application of dynamic compression ………………………….

6.2.3 Cell Proliferation Assay ………………………………………..

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6.2.4 Sulphated glycosaminoglycan (S-GAG) quantification

assays ……………………………………………………………...

6.2.5 Quantitative Polymerase Chain Reaction (q-PCR) …………….

6.3 Results …………………………………………………………………

6.3.1 Cell proliferation and morphology …………………………….

6.3.2 Proteoglycan synthesis …………………………………………

6.3.3 Gene Expression ……………………………………………….

6.4 Discussion ……………………………………………………………..

6.4.1 Cell proliferation and morphology …………………………….

6.4.2 Proteoglycan synthesis …………………………………………

6.4.3 Gene Expression ……………………………………………….

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CHAPTER 7: CONCLUSIONS AND FUTURE DIRECTIONS ………

7.1 Conclusions ………………………………………………………........

7.2 Future directions ……………………………………………………….

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REFERENCES …………………………………...………………………. 116

PUBLICATIONS ……………………...…………………………………..

APPENDIX I ………………………………………………………………

APPENDIX II ……………………………………………………………...

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LIST OF FIGURES

Figure 2.1: Zones of Articular cartilage, modified from (Bhosale & Richardson,

2008…………...………………………………………………………..........................11

Figure 2.2: Wutizite structure of ZnO crystal ………………………………………….32

Figure 2.3: Chondrocytes cultured in a T-75 flask losing their morphology and turning

flat only after 24 hours of culture ………………………………………………………34

Figure 3.1: (A) The reaction scheme of POC pre-polymer fabrication via the

polycondensation reaction in a two necked round bottom flask with a nitrogen purge to

remove excess water content. (B) Synthesis scheme of monomers reacting to form POC

polymer ……………………………………………………..………………………….37

Figure 3.2: Steps involved in fabrication of (A) Pure-POC and (B) POC-ZnONP

scaffolds, films and coatings prior to curing ……………………………....……………39

Figure 3.3: A custom-designed rack placed under water to categorise Pure-POC scaffolds

from ZnO-POC scaffolds and to leach out the salt from scaffolds ……….......…………40

Figure 3.4: Procedure of isolation of primary chondrocytes from cow legs. (A) Removal

of hide from cow leg insert shows completely removed hide, (B) Opening of joint by

removing ligament, (C) Cutting of cartilage and (D) Complete shaved cartilage joint and

explants placed in sterile container ……………………………………………......……41

Figure 4.1: FESEM images of POC (control) and 1%, 3% and 5% ZnO-POC composite

scaffolds; (A) POC, (B) 1% ZnO-POC, (C) 3% ZnO-POC and (D) 5% ZnO-POC; inserts

show representative magnified surfaces of pore walls within scaffolds (inserts: A and B

bars = 1 µm; C and D bars = 2 µm) ……………………………………………...………60

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Figure 4.2: Optical microscopy images of Pure-POC and ZnO-POC surfaces developed

by spin-coating technique at a magnification of 100X; (A) POC, (B) 1% ZnO-POC, (C)

3% ZnO-POC and (D) 5% ZnO-POC …………………………………………..………61

Figure 4.3: Water-in-air contact angle results measured for thin films of Pure-POC and

1%, 3% and 5% ZnO-POC produced by spin-coating technique ……………………….62

Figure 4.4: Percentage swelling in water for Pure-POC (control) and 1%, 3% and 5%

ZnO-POC composite films ……………………………………………………………..63

Figure 4.5: Representative ATR-FTIR spectra of POC and 5% ZnO-POC scaffolds …64

Figure 4.6: Thermal degradation curves of Pure-POC and (1%, 3% & 5%) ZnO-POC

scaffolds measured by TGA ……………………………………………………………65

Figure 4.7: Weight loss of Pure-POC, 1%, 3% and 5% ZnO-POC scaffolds over the

period of 20 weeks in Phosphate Buffered Saline (PBS) ………………………………66

Figure 4.8: In-vitro Release kinetics profile of ZnO (Zn2+) from ZnO-POC scaffolds in

PBS at 37°C; Mt = mass of ZnO released at time intervals; M∞ = total mass of ZnO within

the scaffold (all Mt /M∞ ratios are the mean values for five samples measured with

AAS)………………………………………………………........………………………67

Figure 4.9: Zn-ions released over the period of 28 days calculated in ppm via AAS.

Where * shows a significant difference between the number of days, p

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Figure 4.11: (A) Resazurin reduction (%) for Pure-POC (control) and 1%, 3% and 5%

ZnO-POC scaffolds after 24 and after 72 hours. FESEM images of scaffolds after 72

hours in chondrocyte culture: (B) Pure-POC; and (C) 1% ZnO-POC (bar = 10

µm)……………………………………………………………………….…………….70

Figure 5.1: Summary of experimental design ……………………………...…………81

Figure 5.2: Cell vitality determined via resazurin reduction assay, of chondrocytes,

seeded on Pure-POC and 1% ZnO-POC in normoxic (21%O2) & hypoxic (5%O2)

conditions. Where * shows a significant difference between O2 tension. 5≤ N ≤ 6,

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differences between groups and * shows a significant difference between O2 tension. 5≤

N ≤ 6, p

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Figure 6.5: Confocal imaging of Chondrocytes (A-H) seeded on Pure-POC and 1% ZnO-

POC scaffolds in normoxic (21%O2) & hypoxic (5%O2) conditions, under static and

dynamic culture regimes. Outer images are at 10X with a scale bar of 300 μm and their

respective inserts are at 40X magnification with a scale bar of 100 μm……………….102

Figure 6.6: SEM imaging of Chondrocytes (A-H) seeded on Pure-POC and 1% ZnO-

POC scaffolds in normoxic (21%O2) & hypoxic (5%O2) conditions, under static and

dynamic culture regimes. Outer images are taken at 1000X magnification, it has a scale

bar of 10 μm. Corresponding images in inserts were taken at 3000X magnification having

a scale bar of 2 μm……………………………………………………………………104

Figure 6.7: DMMB assay for glycosaminoglycans (A) Total S-GAG, measured in

digested scaffolds and (B) S-GAG measured in medium via DMMB assay for

chondrocytes seeded on pure-POC and 1% ZnO-POC in normoxic (21%O2) & hypoxic

(5%O2) conditions, under static and dynamic culture regimes. Where ^ shows a

significant difference between due to the presence of ZnONP, * shows a significant

difference between O2 tension and + shows a significant difference between static and

dynamic culture. p

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LIST OF TABLES

Table 2.1: Constituents of articular cartilage with their respective composition……….8

Table 3.1: Ingredients to prepare chondrocyte culture medium…………….…………..42

Table 3.2: Ingredients to prepare papain digest buffer………………….………………45

Table 3.3: Ingredients used in preparation of DMMB solution for DMMB

assay………………………………………………………………….....………………47

Table 3.4: Gene identities and their respective amplicon length from

LifeTechnologies…………………………………….....………………………………50

Table 4.1 Compression properties and porosities of ZnO-POC scaffolds……………....62

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List of Abbreviations

3D three dimensional

AAS atomic emission spectroscopy

ACAN Aggrecan gene

ACI Autologous Chondrocyte Implantation

ANOVA Analysis of Variance

ATCC American Type Culture Collection

BAC Bovine Articular Chondrocytes

BMPs Bone Morphogenic Proteins

CA Citric Acid

cDNA complimentary Deoxyribonucleic acid

CLSM confocal laser scanning microscopy

COL2A1 Collagen Type II – Alpha 1 gene

COLX Collagen Type X gene

COMP Cartilage Oligomeric Protein

DMEM Dulbecco’s Modified Eagle’s Medium

DMMB Dimethyl Methylene Blue

DNA Deoxyribonucleic acid

ECM Extra Cellular Matrix

EDTA Ethylenediaminetetraacetic acid

EDX Energy Dispersive X-ray

em emission

EPS equilibrium percentage swelling

ex excitation

FBS Foetal bovine serum

FDA Food and Drug Administration

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FESEM Field Emission Scanning Electron microscope

FTIR Fourier transform infrared spectroscopy

GAGs Glycosaminoglycans

GAPDH Glyceraldehyde-3-phosphate dehydrogenase gene

GRAS Generally Regarded as Safe

HA Hydroxyapatite

HCl Hydrochloride

HEPES 4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid

HIF-1A Hypoxia Inducible Factor-1 alpha gene

HNMR Hydrogen nuclear magnetic resonance

LB-Broth Luria-Bertani broth

MALDI-TOF Matrix-assisted laser desorption ionization- Time of Flight

min minutes

MMP-13 Matrix Metalloproteinases – 13 gene

MP-AES microwave plasma-atomic emission spectrophotometer

N2 Nitrogen

NaCl Sodium Chloride

NaOH Sodium Hydroxide

NP Nanoparticles

O2 Oxygen

OD 1, 8 Octanediol

PBS Phosphate Buffered Saline

PCL Polycaprolactone

PEG Poly Ethylene Glycol

PGA Poly Glycolic Acid

PGS Poly (glycerol sebacate)

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PLA Poly D, L – Lactic acid

PLGA Poly D, L – Lactic acid-co-glycolic acid

PLLA Poly-L- Lactic acid

POC Poly 1, 8 Octanediol Citrate

PPLG Poly γ-propargyl-l-glutamate

Ppm parts per million

PVP polyvinylpyrolidone

q-PCR Quantitative Polymerase Chain Reaction

RA Rheumatoid Arthritis

RNA ribonucleic acid

ROS reactive oxygen species

SEM Scanning Electron Microscopy

S-GAG Sulphated Glycosaminoglycan

SOX9 SRY-box 9

SRTR Scientific Registry of Transplant Recipients

SSC Saline Sodium Citrate

TCP Tricalcium Phosphate

TE Tissue Engineeirng

TGA thermo gravimetric analysis

UNOS United Network for Organ Sharing

w/v weight by volume ratio

w/w weight by weight ratio

WACA Water-in-air contact angle

ZnO Zinc Oxide

ZnONP Zinc Oxide Nanoparticles

β-TCP Beta Tricalcium Phosphate

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CHAPTER 1: INTRODUCTION

1.1 Thesis Overview

This thesis comprises 7 chapters. Chapter 1 describes the background of the research and

the research problem, followed by objectives and hypothesis. The second chapter

discusses the related literature in depth, and lastly, the motivation to perform the research

is included. A complete methodology and list of materials used is provided in chapter 3

so the reader can repeat the experiment if they want to. Finally, chapter 4, 5 and 6 are

divided according to the objectives, and each chapter includes the results and discussion

for the defined objectives. Chapter 7 is the last chapter in this thesis and provides the

essence of the thesis in the form of a conclusion, and future direction for related research.

1.2 Background

As per 17 May 2012, data provided by Organ Procurement and Transplantation

Network/Scientific Registry of Transplant Recipients (SRTR)/ glyco (UNOS) shows,

there are more than 114,000 people in the queue for organ transplants. Every 10 minutes

a new patient is added to the queue and 18 people die each day due to organ shortages

(Network, 2012). Each year, more than one million people are in need of heart valves,

corneas, skin cardiovascular tissue, and bone tissue (Network, 2012).

The facts and figures mentioned above only represent the United States of America. The

rest of the world population, especially the third world countries where quality of life is

poor and organ donations are scarce, are also affected. There is an urgent need for

biocompatible, efficacious, and cost-effective engineered tissues and organs, and this

need only increases with advances in quality of life and life expectancy.

Articular cartilage is by far one of the most essential and yet functionally complex tissues

that scientists are trying to repair, yet minimal success has been for decades. A major

cause of disability is the joint pain that cartilage disease produces, which is not only

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limited to older people but found in middle aged persons as well. The cause of joint pain

is degeneration of articular cartilage due to primary osteoarthritis or trauma (Temenoff &

Mikos, 2000). Cartilage is a tissue with limited self-repair capabilities, and hence any

kind of injury to the cartilage tissue which is sustained for a substantial time eventually

leads to further detrimental effects (O'Driscoll, 1998; Steinert et al., 2007; Temenoff &

Mikos, 2000). Different approaches are adopted by clinicians to address the issue of

cartilage repair depending upon the nature of cartilage damage. However current

treatment options fail to secure a permanent solution. Furthermore, infection and

inflammation may lead to treatment failing. One of the major causes of biomaterial

infections is peri-operative contamination and research has shown that tissue engineered

implants also pose a risk similar to biomaterial associated infections (Kuijer et al., 2007).

Amongst total knee arthroplasty patients, a risk of peri-operative infection of up to 38%

has been reported recently (Janz et al., 2015). Matrix-induced Autologous Chondrocyte

Implantation has also been reported to cause superficial infection among patients (Bartlett

et al., 2005). Additionally, a recent study demonstrated that only the presence of nano

surface topography can help reduce bacterial adhesion while increasing cell proliferation

due to nanoscale roughness (Liu et al., 2015). This finding adds to the use of

nanomaterials, which are already employed in the delivery of drugs (Pi et al., 2015; Shi

et al., 2015), fighting of infection (Kalashnikova et al., 2015), enhancing mechanical

properties of scaffolds (Pooyan et al., 2015) and increasing cell proliferation (Holmes et

al., 2016).

Chondrocytes are the only cell type that is present in cartilage. These cells flourish in a

hypoxic environment, while the Extra Cellular Matrix (ECM) in which these cells are

embedded undergoes dynamic compression. Current research lack these native

physiologic stimuli and other biofactors as required for chondrocytes to flourish.

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Tissue engineering provides an alternative approach to address existing problems in an

effective and efficient way. Even though cartilage tissue engineering have been adopted

by scientist to cure the dilemma of cartilage damage. However, till date there are many

problems that are associated directly and indirectly that impede implementation of

cartilage repair strategy via tissue engineering approach.

1.3 Problem Statement

Various solutions have been put forward by scientists to successfully implement cartilage

tissue engineering clinically. However, the limitations of their proposed techniques

exceed their benefits. the use of hydrogels, natural polymers, synthetic polymers, and

metal implants has been suggested but they have generally proved unsuitable for clinical

use, due to their inferior mechanical properties (Johnstone et al., 2013; Spiller et al., 2011)

and lack of recovery of the implanted material to its original level due to creep effect (Shi

et al., 2016). Moreover, current cartilage repair strategies have raised a multitude of

concerns, including infection (Lichstein et al., 2015; Yeo et al., 2015) implant failure

(Hazelwood et al., 2015; Vahdati & Wagner, 2013), lack of native tissue architecture

(Chung & Burdick, 2008; Doran, 2015) and absent cell morphology (Hardingham et al.,

2002). In order to fight any bacterial infection many antibacterial agents have been

studied (Aksoy et al., 2010; Davies, 2003). However, Zinc Oxide stands out as favourable

because it also plays a significant role in cartilage synthesis(Kirsch et al., 2000). Current

clinical methods include minimally invasive surgery to resurface articular cartilage,

arthoscopic lavage, or marrow stimulation techniques. Unfortunately all the

aforementioned clinical methods currently used are palliative and for temporary relief

only. The most frequently used treatment options for cartilage tissue repair have

limitations that hinder their long-term clinical implementation (Marlovits et al., 2006;

Portocarrero et al., 2013).

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Furthermore, only a handful of studies are present that implement all three components

of cartilage tissue engineering: (a) cells, (b) scaffold, and (c) bio factors, together. More

recent findings signify the importance of using a three-dimensional environment with

dynamic loading for cartilage tissue engineering (Panadero et al., 2016).

A cartilage tissue engineering approach must benefit from a suitable biocompatible and

biodegradable scaffold that will provide a micro structure for the chondrocytes to settle

and proliferate, while being porous enough to allow cell-cell interaction and exchange of

nutrients for uniform tissue formation. Moreover, the newly fabricated scaffold must also

be flexible enough to be implanted via minimally invasive techniques, where required

(Johnstone et al., 2013).

A successful intervention for a cartilage repair strategy must involve external stimuli that

will deliver a physiologic environment to cultured chondrocytes under a 3D structure.

Furthermore, to tackle inflammation and infection, a suitable anti-inflammatory and

antibacterial mediator must be integrated to assist with diminishing these effects while

having no adverse effects on the cells.

In essence, the combination of biocompatible scaffold, physiologic stimuli and

facilitation of antibacterial and anti-inflammatory properties bears promise for more

effective treatment of cartilage disease.

1.4 Objectives

The objectives of this research are:

1. To develop a biodegradable scaffold with controlled pore size and porosity

mimicking the properties of native articular cartilage.

2. To select and optimise the concentration of nano particles that must be

biocompatible and having antibacterial effect while showing no toxicity towards

chondrocytes.

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3. To study the independent and combined effect of Hypoxia and ZnONP on bovine

articular chondrocytes seeded on polymer scaffolds.

4. To establish a relationship between dynamic mechanical stimuli, ZnONP and

oxygen tension in chondrocyte seeded polymer scaffolds.

1.5 Hypotheses

1. It is hypothesised that chondrocytes will proliferate, secrete more extracellular

matrix, and show higher gene expression under a hypoxic environment.

Furthermore the incorporation of nanoparticles will deliver an anti-bacterial

mechanism to the fabricated scaffold.

2. Furthermore it is hypothesized that scaffolds that will go under dynamic

compression with a combination of hypoxia and nanoparticles, will eventually

display a more chondrogenic morphology and will help in elevating cartilage

specific matrix gene while lowering catabolic gene expression.

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CHAPTER 2: LITERATURE REVIEW

2.1 Introduction

In the present study, a novel composite biomaterial for use within cartilage tissue

engineering is developed, optimised, and evaluated. The literature which underpins this

is therefore reviewed, starting with the current understanding of cartilage as a tissue,

followed by cartilage injury, and a review of cartilage repair options. From here, the focus

moves to the three components of a tissue engineering strategy as applied to cartilage

tissue engineering; first, cell source; second, cell signalling strategy (covering

biochemical and biomechanical signalling); and finally, the scaffold material (including

a focus on the target material for the present study).

2.2 Cartilage

Cartilage is a tissue that lacks blood vessels, nerves and lymph vessels, and is composed

of sparingly distributed chondrocytes (only 1% in humans) (Poole, 1997). There are 3

types of cartilage in the human body, depending upon matrix composition; elastic

cartilage, fibrocartilage and hyaline cartilage (Jung, 2014). Elastic cartilage makes up the

flexible cartilaginous structures of the nose and ear. It contains elastin as an additional

component to ECM. Fibrocartilage contains a higher content of collagen in ECM than in

hyaline cartilage and is present in ligaments and tendons that are located in close

proximity to bone (Temenoff & Mikos, 2000). Hyaline cartilage is mainly found in the

larynx, trachea, ribs, knee and bronchus (Jung, 2014).

Long bones have a covering at their articulating ends known as hyaline cartilage. This

type of cartilage is more commonly referred to as articular cartilage. This hyaline cartilage

provides a low-friction surface with wear resistant properties (Bhosale & Richardson,

2008).

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Hyaline cartilage is made up of water and ECM. Within this ECM, type II collagen is

most abundant, followed by proteoglycans that suffice for mechanical properties in vivo

(Bhosale & Richardson, 2008). For an articular joint to function optimally, it is critical

that engineered cartilage must comply with the properties of native articular cartilage. It

must qualify to withstand load-bearing forces, have low friction a coefficient, and have

other necessary biological and mechanical properties (Doran, 2015).

2.2.1 Composition of articular cartilage

2.2.1.1 Chondrocytes

Among all the tissues of the human body, articular cartilage has the lowest cellular

density, according to volume. Humans have only 1% cellular material in their articular

cartilage, namely cells called chondrocytes (Temenoff & Mikos, 2000). Chondrocytes are

sparse but extremely important for replacement of degraded matrix molecules in order to

maintain homeostasis. At week 5 of gestational, a blastema is formed when few

mesenchyme cells form an aggregate. This blastema starts secreting a cartilage matrix

which forms the foundation of chondrocytes (Bhosale & Richardson, 2008). It is reported

that a cilia extends from surface of some chondrocytes into the ECM that contributes to

modifying ECM properties in response to mechanical stimulus (Buckwalter & Mankin,

1997). Chondrocytes pass through various lineages, producing proteins vital for ECM.

Chondrocytes in the periphery secrete collagen to form a hyaline cartilage. When

chondrocytes mature, they tend to halt their activity, appear circular and are enclosed in

a matrix (Buckwalter & Mankin, 1997; Hall, 1994). Table 2.1 summarises the

constituents of articular cartilage.

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Table 2.1: Constituents of articular cartilage with their respective composition

(Jung, 2014; Poole, 1997; Temenoff & Mikos, 2000).

Constituent Wet weight (%)

Dry weight

(%)

Cellular

Material Chondrocytes ----- 1

Extracellular

Matrix Water 80 -----

Macromolecules Collagen Proteoglycans Non-collagenous

proteins

20-40 10-20 10-20 -----

----- 50-60 25-35 15-20

2.2.1.2 Extracellular Matrix

Apart from chondrocytes, articular cartilage comprises a complex mix of water and

structural macromolecules in a framework that provides stability to the articular cartilage.

A combination of water and macromolecules is called ECM. The majority constituent in

ECM of articular cartilage is water, approximately 80% of the weight. However in

different regions of cartilage the percentage of water varies from 60% to 80%. The

macromolecules comprise collagens, proteoglycans, and non-collagenous proteins.

Macromolecules (when combined) account for 20 – 40 % wet weight of articular cartilage

(Jung, 2014).

2.2.1.3 Water

In articular cartilage water accounts for majority of the wet weight of the tissue; around

80%. Apart from water tissue fluid contains bulk of cations to neutralise negatively

charged GAG’s in ECM. Tissue fluid also contains metabolites and gases. The presence

of tissue fluid is of vital importance to avascular cartilage as it helps the exchange of

nutrients and oxygen (O2) with synovial fluid. Furthermore, recovery of cartilage and

resistance to compression is achieved via the presence of tissue fluid in ECM (Buckwalter

& Mankin, 1998b).

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2.2.1.4 Collagens

The second-most abundant constituent of ECM is collagen, which accounts for 10% to

20% of the wet weight of articular cartilage (Bhosale & Richardson, 2008) while it

equates to 50% to 60% of the dry weight of articular cartilage (Jung, 2014). Collagen

types II, VI, IX, X and XI are found in articular cartilage. However, the major role player

is type II collagen, which represents 90 – 95% of all the collagens present. Collagen type

II also forms the foundation of interlinked fibrils in articular cartilage (Buckwalter &

Mankin, 1997). The interlinked fibrils formed by type II collagen not only provide

physical strength to the cartilage but also trap other macromolecules (Cohen et al., 1998;

Temenoff & Mikos, 2000).

2.2.1.5 Proteoglycans

The third constituent of ECM are the proteoglycans, proteins that give compressive

capabilities to the cartilage. However they are only 10 – 20 % of the total wet weight

(Bhosale & Richardson, 2008), and they account for 25% to 35% in dry form (Buckwalter

& Mankin, 1997). The proteoglycans are composed of polysaccharides and proteins in

the ratio of 19 to 1 respectively (Wirth & Rudert, 1996). The protein forms the core, with

attachment to one or more glycosaminoglycans (GAGs) chains (Buckwalter & Mankin,

1997; Temenoff & Mikos, 2000). GAGs are from a family of polysaccharides which have

a repeating disaccharide unit (Afratis et al., 2012; Anower-E-Khuda & Kimata, 2015).

Chondroitin sulfate, dermatan sulfate, keratin sulfate and hyaluronic acid are all GAGs

that are present in articular cartilage (Brézillon et al., 2014; Hall, 2012). Proteoglycans

are sub-divided into 2 groups: (a) large aggregating proteoglycans and (b) small

aggregating proteoglycans. Large aggregating proteoglycans are more commonly

referred as Aggrecan (Buckwalter & Mankin, 1997). The Aggrecans are responsible for

resilience and distribution of stress in articular cartilage (Temenoff & Mikos, 2000).

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2.2.1.6 Non-Collagenous proteins and Glycoproteins

Non-Collagenous proteins and Glycoproteins account for about 15% to 20% of articular

cartilage in dry weight. In an articular cartilage there exists wide variety of these

molecules. They possess minute quantities of oligosaccharide attached to a protein core.

Contrary to proteoglycans and collagens, these non-collagenous proteins and

glycoproteins have not been studied extensively. However, they tend to help in

maintenance and organization of macromolecular network of ECM(Buckwalter et al.,

2005; Gallagher et al., 1983). Among them, cartilage oligomeric protein (COMP) is most

abundant. It is concentrated in the matrix of chondrocyte cells and has a binding affinity

chondrocytes (Buckwalter & Mankin, 1997; Temenoff & Mikos, 2000); COMP is

involved in maintaining the chondrogenic phenotype (Aigner & Fan, 2003).

2.2.2 Zonal arrangement in articular cartilage

Articular cartilage is divided in to 4 zones (figure 2.1) according to their structure and

function.

1. Superficial zone

2. Transitional zone

3. Deep zone

4. Calcified zone

1. Superficial zone

The superficial zone comprises of flat cells and is the thinnest layer. It is laminated by a

specialised layer of synovial fluid called the lamina splendens (Instructional Course

Lectures, The American Academy of Orthopaedic Surgeons - Articular Cartilage. Part I:

Tissue Design and Chondrocyte-Matrix Interactions*†, 1997). The lamina splendens is a

protein that enhances the low – friction properties of the surface of articular cartilage

(Buckwalter & Mankin, 1997; Temenoff & Mikos, 2000).

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2. Transitional zone

This zone is dominated by spheroid-shaped cells, although the cell density is lower.

Collagen fibres with a higher diameter can be found in this zone, which also has a high

content of the proteoglycan aggrecan (Bhosale & Richardson, 2008; Buckwalter &

Mankin, 1997).

3. Deep zone

This zone has the lowest of cells densities, with spherical cells distributed throughout the

zone. The proteoglycan concentration is the highest in this region (Bhosale & Richardson,

2008; Buckwalter & Mankin, 1997).

4. Calcified zone

Mineralisation is higher in this region of articular cartilage, meaning that cells are

embedded in a matrix. The cells in calcified zone synthesise Type X collagen, which is

mainly responsible for providing mechanical stability to cartilage, and also assists in

absorbing shocks(Buckwalter & Mankin, 1997; Temenoff & Mikos, 2000).

Figure 2.1: Zones of Articular cartilage, modified from (Bhosale & Richardson,

2008).

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2.3 Chondrogenesis and chondrocyte differentiation

The process by which cartilage forms is called chondrogenesis. This is a natural process

and chondroprogenitor cells are responsible for its initiation. Chondroprogneitor cells are

basically mesenchymal cells. Mesenchymal cells come together and form an aggregate

and start condensing in one area; this marks the beginning of chondrogenesis. Bone

Morphogenic Proteins (BMPs) are responsible for contributing to initiating

chondrogenisis by helping in aggregation of loose mesenchymal cells. In the early stages

of chondrogenesis, the mesenchymal cells secrete ECM and cell adhesion molecules

(Goldring et al., 2006). Among the secreted ECM there are large amounts of type II

collagen (Ng et al., 1993), N-cadherin (Oberlender & Tuan, 1994), N-cam (Widelitz et

al., 1993) and tenascin (Mackie et al., 1987). Apart from the ECM and cell adhesion

molecules, SRY-box 9 (SOX9); a transcription factor, is highly expressed. SOX9 plays a

vital role in chondrogenesis and differentiation of chondrocytes.

Differentiation of mesenchymal cells into chondrocytes causes the cells to secrete the

ECM that contains large quantities of type II collagen and Aggrecan. After early

chondrocyte differentiation, this is the next development stage of cartilage. In this stage

the cells proliferate rapidly providing a template for the cartilage. After a certain time old

cells draw away from the cell proliferation cycle and enter in to a phase of hypertrophic

differentiation. In this process the chondrocytes grow in size, terminally differentiate,

mineralize, and eventually undergo apoptosis (Freed et al., 1998; Zuscik et al., 2008).

After the death of chondrocytes, they degrade and leave a matrix of their cells that serves

as a scaffold for deposition of minerals and formation of bone (Zuscik et al., 2008).

2.4 Articular Cartilage Disease and Injury

Articular cartilage diseases represent the most common cause of joint pain in middle aged

and older people (Buckwalter & Mankin, 1998b), and hospital treatment cost for total

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knee replacement was reported to be US$ 28.5 billion in 2009 in the US alone (Murphy

& Helmick, 2012). Together with acute injury, such cartilage disorders also result in a

substantial reduction in quality of life (Minas, 1999; Murray et al., 2015).

Depending upon the nature of injury, cartilage injury can be subdivided in to three major

categories; Full thickness defect, partial defect and matrix disruption. A full thickness

defect is the most painful and it usually affects entire cartilage thickness while also

penetrating deep into the subchondral bone (Temenoff & Mikos, 2000). In this type of

injury fibrin clot formation occurs at the site of defect. Blood and bone marrow cells bind

to this site and form patches of fibrin clots (Aigner & Fan, 2003). The newly formed

tissue lacks the strength of native articular cartilage and has a higher permeability. Lack

of strength and excess permeability eventually contribute to degradation and permanent

damage to articular cartilage and subchondral bone. An injury of as little as 3mm depth

is considered to be a full thickness defect. (Coutts et al., 1997; Fortier et al., 2011;

Temenoff & Mikos, 2000).

A secondary injury is a partial thickness defect, resulting in disruption to the surface of

cartilage. This damage, unlike a full thickness defect, does not extend to the subchondral

bone. In this type of injury the surrounding cells do proliferate but not to an extent which

can repair the defect completely (Temenoff & Mikos, 2000). Research over the years has

provided evidence against the healing of tissue (Ghadially et al., 1977) or progressive

degeneration (Grande et al., 1989) in the case of partial thickness defects. Not only is

regeneration limited at the cellular level of cartilage, but also negative charges amongst

proteoglycans during injury contribute to the lack of cellular adhesion (Hendrich et al.,

2003).

A blunt trauma to cartilage causes matrix disruption injury. This type of injury may occur

from a sporting activity or accident. If the extent of injury is not extreme the damage is

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reversed by the viable chondrocytes present at the site of injury (Temenoff & Mikos,

2000).

Thus, despite it being such a serious and significant issue, treatment of cartilage problems

is severely restricted due to the limited ability of cartilage to self-repair (Lam et al., 2014).

This often means that degeneration occurs at a higher rate than repair, causing increasing

failure of the degenerative cartilage surface (O'Driscoll, 1998).

2.5 Treatment Options for Cartilage Injury

To provide a sustainable solution to the problem of cartilage damage the primary

objective of every clinician is to yield normal function of the joint(Temenoff & Mikos,

2000). Generally there exists three major interventions to restore joint functionality.

2.5.1 Marrow Stimulation Technique

The marrow stimulation technique involves micro-fracture, abrasion arthroplasty, or

drilling, and is usually adopted as a primary treatment for cartilage defects up to 2 to 4

cm2. This technique involves creating defects that extend to the sub-chondral bone that

eventually result in clotting at the site of defect and a natural repair response being

induced. This response leads to the formation of fibrocartilage repair tissue.

Fibrocartilage tissue is undesirable as it is less resilient than articular cartilage, and

mechanically speaking the properties fail to match that of the native cartilage (Buckwalter

& Mankin, 1998a; Caplan et al., 1997). Such marrow stimulation techniques are

considered to be effective only in younger patients without any prior surgical intervention

(Demange et al., 2014; Mithoefer et al., 2009), and the size of the defect that can be treated

by marrow stimulation technique is limited to 2 to 4 cm2. Further, this technique is

ineffective for the femoral condyle region (Gudas et al., 2005; Kreuz et al., 2006).

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2.5.2 Osteochondral Grafting

Tissue transplantation of periosteum, perichondrium or osteochondral grafts is used for

regeneration of hyaline cartilage (Bouwmeester et al., 1997; Buckwalter & Mankin,

1998a; Hunziker, 2002). The most common approach is to harvest osteochondral

cylinders from the donor site and then fit them into holes already drilled in areas of defect.

The drilled areas are filled via one or more grafts that range from 5 to 11 mm in diameter

while their lengths are 13 to 15 mm (Hangody, 1997). Studies have shown mixed results

in cases of osteochondral transplants. Reduced pain and improved function was reported

after 6.5 years by Outerbridge et al. (Outerbridge et al., 1995). In two separate studies by

Jakob et al. (Jakob et al., 2002), and Hangody et al. (Hangody et al., 2001) improved knee

function and good results were reported in more than 90% of patients. In contrast, Laprell

and Petersen reported a normal knee in fewer than 50% of patients after a follow up of 6

to 12 years (Laprell & Petersen, 2001). A major limitation of this approach is the

availability of tissue where there are large cartilage defects. Further, long term follow up

is missing and this technique is reported to be most appropriate for patients less than 35

years of age (Buckwalter & Mankin, 1998a; Caplan et al., 1997; Hunziker, 2002; Insall,

2001). In addition, osteochondral transplantation is associated with some post-operative

complications. The accumulation of water (joint effusion), bleeding in joint spaces

(hemarthrosis), and persistent swelling have all been reported after osteochondral

transplantation (Bobic, 1999; Bös et al., 2000; Hangody et al., 2001; Hunziker, 2002).

Researchers have also reported loose bodies, donor site pain, and avascular necrosis as

other post-operative complications that emanate due to osteochondral transplantation

(Aigner & Fan, 2003).

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2.5.3 Autologous Chondrocyte Transplantation

Due to the failure of osteochondral transplantation, autologous chondrocyte

transplantation was developed and used since 1994 by Brittberg and coworkers (Brittberg

et al., 1994). Autologous chondrocyte implantation (ACI) is considered to be used as the

last line of defence when other techniques fail. ACI is considered able to repair full-

thickness lesions in cartilage up to 16 cm2. This technique involves isolation of a patient’s

own cells grown in-vitro then implanted at the site of the defect. The follow up results of

ACI after 2 to 9 years have reported a failure in more than 30% patients after treatment

of multiple lesions. Furthermore, patients with patella lesion treatment demonstrated

more than 30% failure, but isolated cartilage lesion treatment displayed less than 10%

failure (Peterson et al., 2000). Another long-term follow up study by Harris J. D. et al.

sought to determine if current literature supports ACI over other interventions. They

observed that ACI can only be helpful over other interventions if the size of defect is

greater than 4 cm2. ACI or micro fracture or osteochondral transplants, in summary, only

provide short term success.(Harris et al., 2010) A study by Niemeyer P and co-workers

reported that complications like graft failure, delamination, chondromalacia (softening of

cartilage) and hypertrophy can arise after ACI (Niemeyer et al., 2008). More recent

reports by Harris J. D et al. have reported decreased failure rates (up to 7.5 %) for ACI,

however there are other reports that continue to demonstrate inconsistent outcomes of

ACI without any clear recommendation. Furthermore, it was also reported that the

differences between various treatments were so small that their clinical significance is

questionable, beyond mere statistical significance (Vavken & Samartzis, 2010). Lately it

was reported that ACI improves movement significantly after at least 24 months (von

Keudell et al., 2016).

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2.6 Cartilage Tissue Engineering

Current treatment options, like ACI, Osteochondral transplants and marrow stimulation

have demonstrated variable success rates, however more long-term results are not

satisfactory (Kreuz et al., 2006; Niemeyer et al., 2008; Redman et al., 2005; Vavken &

Samartzis, 2010). The major limitation that arises due to the use of the aforementioned

therapeutic techniques for cartilage repair is the lack of mechanical properties of newly

formed tissue compared to native articular cartilage tissue. Due to these inferior

mechanical properties, the newly formed tissue is prone to failure (Hunziker, 2009).

Cartilage tissue engineering strategies contribute to a durable and functional replacement

option, beyond any other surgical intervention or technique (Kock et al., 2012).

Tissue engineering is a 3 tier based approach, involving the combination of living cells,

a cell signalling strategy, and a scaffold (Hendrich et al., 2003). Cartilage tissue

engineering allows researchers a great deal of flexibility in combining various techniques

independently and in combination to address the issue of cartilage repair. Cartilage tissue

engineering also allows the selection of various cell types to be implemented in-vitro for

a better understanding of cartilage repair. Moreover, researchers have shown positive

effect of signalling molecules in enhancing anabolic effects and increased ECM synthesis

(Elder & Athanasiou, 2009; Seifarth et al., 2009). Further, a widely researched area for

better fabrication of cartilage tissue is external mechanical stimuli. Direct dynamic

compression has been demonstrated to increase ECM production while improving the

compressive properties of the engineered tissue (Bian et al., 2010; Chen et al., 2015; Luo

et al., 2015). Another leverage given by a cartilage tissue engineering strategy to the

researchers is modification of the properties of scaffolds. Scaffolds provide a temporary

frame for the cells to attach to and proliferate while the tissue is being formed and the

scaffold is being degraded. Recent publications have shown enhanced mechanical and

biological properties of newly fabricated tissue using polymeric scaffolds (Camarero-

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Espinosa et al., 2016; Schulze-Tanzil, 2015). Various cartilage tissue engineering studies

have been undertaken to engineer a tissue with native sGAG content. However, they have

failed to achieve the collagen content present in native articular cartilage, and which gives

the cartilage its tensile properties (Eyrich et al., 2007; Kock et al., 2012). Though there

are many advantages to cartilage tissue engineering, nevertheless finding optimal cell

source is the first issue, followed by choice of scaffold material, the role and the effect of

biochemical and mechanical stimulus and lastly tackling any infection that might occur

due to surgical intervention, amongst others are pitfalls that need to be addressed

(Bhattacharjee et al., 2015; Kock et al., 2012). Gaut C. and Sugaya K. recently suggested

that in order to obtain a clinically effective solution for cartilage tissue engineering; a

combination of factors is required, such as; cell signalling, scaffold material, mechanical

and biological stimulus must be considered (Gaut & Sugaya, 2015).

Given the clear advantages of Cartilage Tissue Engineering, but yet the continued

inability to fully restore cartilage functionality in the long term, the present study seeks

to further advance this technique toward an effective and lasting repair.

2.6.1 Cell Sources

A source of cells is one of the key components of a Tissue Engineering strategy. There is

no current consensus on the optimal source of cell for cartilage tissue engineering, with

division between the use of stem cells, fibroblasts, or chondrocytes – each with their

relative merits and disadvantages.

Stem cells are particularly attractive due to their differentiation potential and the fact they

may be harvested from various tissues. Though stem cells overcome the problem of

limited supply of cell as they can be easily harvested from the patient’s body itself (from

bone marrow (Boeuf & Richter, 2010) and adipose tissue (Gir et al., 2012)), bone marrow

stem cells have lower mechanical properties and matrix synthesis when compared to

chondrocyte seeded scaffolds (Thorpe et al., 2010) and adipose stem cells demonstrate a

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lesser potential than bone marrow stem cells despite being capable of being differentiated

into chondrocytes (Kock et al., 2012). Furthermore, embryonic stem cells tend to

differentiate hypertrophically and moreover, there exists many ethical and legal issues to

overcome before considering embryonic stem cell use (Jukes et al., 2008). Mesenchymal

stem cells are another potential source, however they have demonstrated high cartilage

hypertrophy markers such as Collagen type X (COL X) & Matrix (MMP-13 (Mueller &

Tuan, 2008).

Chondrocytes are well suited to cartilage tissue engineering because they are the native

cells of cartilage tissue (Nicoll et al., 2001). However, they have limited availability and

generally need to be harvested from a monolayer to get an high number of cells, meaning

they eventually lose their cell phenotype (Cournil-Henrionnet et al., 2008).

Another potential source of cells for cartilage tissue engineering are the articular cartilage

progenitor cells that are useful in producing the amount required for cell culture but do

not have the capability to produce cartilage matrix (Zhou et al., 2014). It is reported that

they may also cause donor site morbidity (Mathur et al., 2012).

A new approach should be using chondrocytes in an in-vitro culture systems that mimics

the native environment to overcome the limitation of limited cell numbers and allows

maintenance of cell phenotype.

2.6.2 Cell Signalling Strategies

A key aspect of Tissue Engineering is the integration of a cell signalling strategy

alongside the cellular and scaffold components. This may include aspects such as

biochemical cues, facilitation or stimulation of cell-cell interactions, biomechanical

signalling, or manipulation of the environmental factors involved.

Human body tissues have various associated stimuli that help regulate normal function of

the tissue. Whether it is electrical, mechanical, or chemical, body tissues undergo

different kinds of stimulus; heart tissues undergo electrical stimulation (Cannizzaro et al.,

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2007; Godier-Furnémont et al., 2015), bone tissues are subjected to tension and

compression (Nyman et al., 2009; Romanos et al., 2015), and similarly, cartilage tissue

undergoes compression, hydrostatic pressure, and shear strain (Ryan et al., 2009; Schätti

et al., 2015). Cartilage tissue also experiences a unique stimulus, which is tension in

oxygen (Portron et al., 2015). This is mainly due to the avascular nature of cartilage tissue

(Hansen et al., 2001). When considering cartilage tissue, in particular, successful tissue

engineering will be incomplete without a mechanical stimulus and low oxygen tension.

2.6.2.1 Mechanical stimulus

In the last decade, researchers have increasingly realised the importance of external

mechanical stimuli to enhance the quality of produced cartilage (Temenoff & Mikos,

2000). However, much conflicting information has been reported regarding how

chondrocytes behave in response to mechanical stimulation (Lee et al., 2005).

Physiologically, articular cartilage primarily undergoes compression at joint level (Lee et

al., 2005). In an average human pressure of up to 1MPa during standing, 0.4MPa while

walking, and up to 20MPa while standing up from a chair have been reported (Gooch &

Tennant, 1997; Lee et al., 2005; Urban, 1994). A range of factors are affected by the

compressive loading of articular cartilage. These include biochemical and physical

gradients of nutrients, ion concentrations, electrical charge, and pH (Gray et al., 1988;

Guilak et al., 1995). Different studies have shown that a mechanical stimulus positively

influences cartilage development (Heath & Magari, 1996). The way ECM is composed

and organised provides the cartilage with its biomechanical properties. The presence of

type II collagen helps strengthen the cartilage while proteoglycans build up compressive

resistance in the cartilage (McMahon et al., 2008). Grodzinsky A.J. and co-workers

reported that mechanical stimuli on chondrocytes is essential in order to maintain the

integrity of cartilage (Grodzinsky et al., 2000).

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In its native environment cartilage is usually compressed to about 10% (Sanchez-Adams

et al., 2014). Various studies showcase the importance of dynamic compression.

Researchers have studied the effects of compressing the cartilage from about 5% to 50%

with frequencies ranging from 0.001 Hz to 1Hz using various polymers and gel constructs

(Albro et al., 2013; Appelman et al., 2011; Chen et al., 2015; Chowdhury et al., 2006;

Chowdhury et al., 2003; Davisson, Twana et al., 2002; Démarteau et al., 2003; Di

Federico et al., 2014; Hunter et al., 2002; Jeon et al., 2012; Kisiday et al., 2004; Li et al.,

2013; Morel et al., 2005; Piscoya et al., 2005). The majority of the researchers have

favoured dynamic compression and have demonstrated a higher GAG content with better

cell proliferation and increased cartilage matrix specific gene expression of Type II

collage (COL2A1) and Aggrecan (ACAN) (Appelman et al., 2011; Chen et al., 2015;

Davisson, Twana et al., 2002; Jeon et al., 2012; Kisiday et al., 2004). However,

contradicting results are present for the amount of compression. Some studies support

dynamic compression in the range of 0.001% to 31% (Appelman et al., 2011; Chen et al.,

2015; Davisson, Twana et al., 2002; Démarteau et al., 2003; Jeon et al., 2012) while others

have reported detrimental effect of dynamic compression ranging from 15% and up to

50% (Davisson, Twana et al., 2002; Di Federico et al., 2014; Hunter et al., 2002; Li et al.,

2013; Piscoya et al., 2005). From the current literature it is evident that a dynamic

compression beyond physiologic levels of 10% has proved to be detrimental for cartilage.

One thing is noteworthy here: most of the studies executed to study the behaviour of

cartilage tissue or cartilage cells are performed on agarose, hydrogels, alginate and type I

collagen gels. All the aforementioned constructs/ gels can only be used in laboratory

settings and lack the physical capabilities to be implanted clinically. Limited groups have

reported dynamic compression on polymer/ polymer composite scaffolds which they

deem to be physically compatible. Another important factor that seems neglected in the

current study is the physiologic oxygen level. The majority of the studies of dynamic

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compression in the domain of cartilage tissue engineering were carried out under normal

oxygen settings without comparing the effect of atmospheric and physiologic oxygen

concentration on dynamic compression.

2.6.2.2 Oxygen Tension/Hypoxia

Research suggests that the nature of articular cartilage is avascular; hence it receives its

required nutrients and oxygen by synovial fluid following a passive diffusion mechanism

(Malda et al., 2003). In relation to rest of the body tissues, articular cartilage experiences

oxygen concentrations in the range of 1% to 10% as compared to 21% for other tissues

(Co et al., 2014). Chondrocytes are reported to thrive in oxygen concentration range of

5% to 10% (Grimshaw & Mason, 2000), while losing their activity in anoxic conditions

because 1% of Oxygen concentration is reported to exist in an abnormal or due to any

pathological condition (Grimshaw & Mason, 2000; Zhou et al., 2004). Hypoxia Inducible

Factors are responsible for chondrocytes adapting to low oxygen tension (Fedele et al.,

2002). Various researchers have reported that hypoxia in chondrocytes has resulted in

increased synthesis of ECM proteins in in-vitro conditions (Domm et al., 2002; Madeira

et al., 2015; Yodmuang, Marolt, et al., 2015). Inhibition of collagen type X had also been

demonstrated by inducing hypoxic conditions. The Collagen type X is major marker of

hypertrophy in chondrocytes during the chondrogenesis of adipose-derived mesenchymal

stem cells (Betre et al., 2006; Portron et al., 2015) and of epiphyseal chondrocytes (Chen,

X. C. et al., 2006). A transcription factor, Hypoxia inducible factor 1 (HIF-1) is mainly

responsible for regulation of hypoxic response in the cartilage tissue (Schipani et al.,

2001). HIF-1α is said to be responsible for angiogenesis and glycolysis while at the same

time causing cells to proliferate (Denko et al., 2003; Goda et al., 2003). Hypoxia further

plays a role in formation of collagen fibrils with the help of transcription factor HIF-1 α

(Takahashi et al., 2000). Current research has reported an individual role of hypoxia for

better engineered cartilage, however, not enough details are present in the current

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literature as to how hypoxia will behave in combination with other bio-factors for

cartilage tissue engineering.

However, many studies have highlighted the importance of hypoxia as a success for tissue

engineering of cartilage (Meretoja et al., 2013; Munir et al., 2014; Schipani et al., 2001).

Researchers have shown that hypoxia induces ECM synthesis in in-vitro chondrocyte

culture (Domm et al., 2002) and also helps chondrocytes to proliferate (Choi et al., 2014).

Another important research by Schipani E. et al. demonstrated that lack of hypoxia causes

chondrocytes to terminally differentiate (Schipani et al., 2001). A study by Hansen U. et

al. reported a combination of hypoxia and hydrostatic pressure to enhance collagen type

II production while delaying expression of Collagen type I for in-vitro cartilage tissue

engineering (Hansen et al., 2001). Further research has also displayed the potential of low

oxygen tension to enhance Aggrecan and Collagen type II gene expression while

decreasing Matrix Metalloproteinase – 13 (a pro catabolic gene) expression (Parker et al.,

2013). Most of the studies were executed on chondrocytes on Tri Calcium Phosphate

(TCP) plates (Hansen et al., 2001), under agarose gel (Chowdhury et al., 2006; Parker et

al., 2013) or Alginate (Jeon et al., 2012). Others have used PEG (Appelman et al., 2011),

PGA (Davisson, Twana et al., 2002), or articular cartilage explants (Piscoya et al., 2005).

However, the above mentioned studies did not used scaffolds able to be used clinically as

they either lack the requisite mechanical properties or do not mimicked native cartilage

structure in-vivo.

2.6.3 Tissue Engineering Scaffolds

Different natural and synthetic materials have been tested as scaffolds. Scaffolds have

been developed by natural polymers (Van de Putte & Urist, 1965; Vandeputte & Urist,

1965), synthetic polymers (Elisseeff et al., 1999; Meinig et al., 1997), composites (Peter

et al., 1998; Zhang & Ma, 1999) and ceramics (Friedman et al., 1998; Hollinger &

Battistone, 1986). Synthetic scaffolds have copious benefits such as a greater availability,

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decreased risk of disease transfer, fabrication as per requirement, and the ability to be

shaped into various shapes and sizes (Jeuken et al., 2016).

Just saying that a material should be biocompatible does not solve all the complications.

A tissue engineered scaffold should be readily available to the surgeon; it should promote

cell ingrowth as well as it must degrade in correct ratio with formation of new tissue

(Holmes, 1979; Levine et al., 1997). Also it should be able to fit into irregular places and

must be soft enough to mould and hard enough to withstand the mechanical loads in daily

routine activities (Lichte et al., 2011), meaning that the scaffold should possess

mechanical strengths of site where it needs to be implanted. The architectural properties

of the scaffold; which includes pore size and porosity, should be in relation to the actual

tissue (Peter et al., 1998). The scaffold should also promote cell adhesion and

proliferation (Lichte et al., 2011). Additionally the design of a scaffold should be

biomimetic, implicating it must imitate all the characteristics of the tissue where it is

going to be implanted; whether they are physical chemical or biological properties.

For the aforementioned reasons the polymeric scaffolds are a promising new approach

towards tissue engineering for repairing cartilage defects. However, use of polymeric

materials alone or with composites as cartilage tissue engineering scaffolds remains

largely unexplored.

Various materials have showed their ability to be used as engineered tissues; this includes

metals, ceramics, polymers and their combinations (Burg et al., 2000; Hidalgo-Bastida et

al., 2007; Hollinger & Battistone, 1986; Liu & Ma, 2004). Among these only using

materials and ceramics independently results in lack of degradability and processability.

On the contrary, if polymers are used with metals and ceramics they allow tailored

properties and proper biodegradability. Various scaffolds based on polymers and polymer

composites are being developed that demonstrate a promising future (Burg et al., 2000;

Liu & Ma, 2004; Peter et al., 1998).

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Particularly for cartilage tissue engineering, various polymers and polymer composites

have been used. Polymers such as Poly (ε-caprolactone) (Li, W.-J. et al., 2005), poly (D,

L – Lactic acid-co-glycolic acid) [PLGA] and PLGA with Hyaluronic based polymers

(Yoo et al., 2005), polyurethane based polymers (Grad et al., 2003), poly (ethylene

glycol)-terephthalate and poly (butylene terephthalate) based polymers (Woodfield et al.,

2004), poly (D, L – Lactic acid) [PLA] based polymers (Camarero-Espinosa et al., 2016),

collagen/PLA and Chitosan/PLA polymers (Haaparanta et al., 2014), poly(γ-propargyl-l-

glutamate) (PPLG) polymers (Ren et al., 2015), Poly (glycerol sebacate) [PGS]

(Kemppainen & Hollister, 2010), and Polyoctanediol citrate polymers (Jeong et al.,

2011).

Among several polymers and polymer composites, Poly 1, 8 Octanediol Citrate (POC)

stands as a potential candidate due to its mechanical and chemical properties that suffice

for cartilage tissue engineering applications. POC has proved to enhance cell attachment

and helps in proliferation of cells, furthermore POC has been shown to enhance matrix

production when seeded with chondrocytes (Kang et al., 2006). Studies have also reported

POC polymer to be more effective than other polymers (PGS & polycaprolactone (PCL))

in enhancing chondrogenesis (Jeong & Hollister, 2010). Moreover, its physical properties

can be tailored to those of native articular cartilage (Kang et al., 2006; Moutos et al.,

2007).

2.7 Poly Octanediol Citrate (POC)

2.7.1 POC, its structure, synthesis, and form

Po